Method and system for photoselective vaporization of the prostate, and other tissue

ABSTRACT

A method for photoselective vaporization of prostate tissue includes delivering laser radiation to the treatment area on the tissue, via an optical fiber for example, wherein the laser radiation has a wavelength and irradiance in the treatment area on the surface of the tissue sufficient because vaporization of a substantially greater volume of tissue than a volume of residual coagulated tissue caused by the laser radiation. The laser radiation is generated using a neodymium doped solid-state laser, including optics producing a second or higher harmonic output with greater than 60 watts average output power. The delivered laser radiation has a wavelength for example in a range of about 200 nm to about 650 nm, and has an average irradiance in the treatment area greater than about 10 kilowatts/cm 2 , in a spot size of at least 0.05 mm 2 .

RELATED AND CONTINUING APPLICATION INFORMATION

This application is a continuation of U.S. patent application Ser. No.10/278,723 filed on 23 Oct. 2002.

Application Ser. No. 10/278,723 is a continuation-in-part of U.S. patentapplication Ser. No. 09/737,721 (now U.S. Pat. No. 6,554,824).

Application Ser. No. 10/278,723 also claims the benefit of U.S.Provisional Application No. 60/336,481, filed 24 Oct. 2001.

Application Ser. No. 10/278,723 also claims the benefit of U.S.Provisional Application No. 60/338,728, filed 5 Nov. 2001; and

Application Ser. No. 10/278,723 also claims the benefit of U.S.Provisional Application No. 60/337,810, filed 5 Nov. 2001.

The present application is related to co-pending U.S. patent applicationSer. No. 10/279,087, filed on 23 Oct. 2002.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention relates generally to laser treatment of softtissue, and more particularly to photoselective vaporization of theprostate PVP, and to photoselective vaporization of other tissue.

2. Description of Related Art

Benign Prostatic Hyperplasia (BPH) is a condition wherein continuedgrowth of the prostate restricts the passage of urine through the lowerportion of the bladder and the urethra. BPH is often treated bysurgically removing excess prostate tissue from the transitional zone ofthe prostate that is pressing on the urethra, which usually relieves thebladder outlet obstruction and incomplete emptying of the bladder causedby the BPH.

Recently, the most commonly employed procedure for removal of excessprostate tissue has been transurethral resection of the prostate, alsoknown as TURP. In the TURP procedure, the surgeon utilizes a standardelectrical cutting loop to shave off small pieces of the targeted tissuefrom the interior of the prostate. At the end of the operation, piecesof excised prostate tissue are flushed out of the bladder using anirrigant.

While effective, the TURP procedure is known to cause numerous sideeffects, including incontinence, impotence, retrograde ejaculation,prolonged bleeding and TUR syndrome. Recently, alternative procedureshave been developed which reduce or avoid the side effects associatedwith TURP. One class of procedures involves “cooking” prostate tissue byheating it to a to a temperature above 45 degrees Celsius, causingtissue coagulation. Typically this is accomplished using electricallyresistive elements such as: radio frequency (RF), microwave, orlong-wavelength lasers. An example of a procedure of this nature isdiscussed in U.S. Pat. No. 6,064,914 by Trachtenberg (“ThermotherapyMethod”). Because these procedures leave the thermally-treated tissue inplace, post-procedure edema, dysuria, and retention rates are relativelyhigh. Further, use of thermal procedures requires the patient to becatheterized for several days following the procedure, and may causeextensive and unpredictable scarring of the intra prostatic urethra.

Another class of procedures involves vaporizing or ablating the targetedtissue using laser light. These procedures generally avoid the highinfection rates and scarring problems of thermally-based procedures.However, laser ablation of prostate tissue has to date, required the useof an expensive laser capable of generating high-power laser light. Thehigh cost of purchasing or leasing such a laser results in a concomitantincrease in the cost of the procedure. Finally, the ablation processtypically occurs slowly, resulting in a lengthy procedure time.

The Ho:YAG laser and its fiberoptic delivery system is an example of alaser that is commonly used for ablating prostate tissue. The Ho:YAGlaser generates pulses of 2100 nm light that are strongly absorbed bywater in the prostate tissue and in the saline irrigant positionedbetween the distal end if the fiberoptic and the tissue. The absorptioncoefficient of water is so high at 2100 nm that 50% of the light isabsorbed within 0.2 mm of water. Consequently even a thin layer ofirrigant positioned between the distal end on the fiberoptic and thetissue will absorb a large fraction of the laser light. Furthermore withthe short pulse durations (Tp<0.5 ms) and large pulse energies (Ep>1.0joule) used for ablating prostate tissue the irrigant is explosivelyboiled creating a shock wave that tears tissue. Because water is such alarge constituent of prostate tissue and blood, there is essentially noselective absorption by blood. This combination of violent tissuedisruption and the superficial unselective light penetration leads topoor hemostasis.

Nd:YAG lasers operating at 1064 nm have also been used for ablatingprostate tissue. Although 1064 nm light is hemostatic at high powerlevels its low absorption in blood and prostate tissue leads toinefficient ablation and a large residual layer of thermally denaturedtissue several millimeters thick. After surgery, the coagulated,thermally denatured tissue swells and leads to transient urinaryretention, which can cause long catheterization times, painfulurination, and high infection rates.

Frequency doubled Nd:YAG lasers operating at 532 nm in a Quasicontinuous mode at power levels up to 60 watts have been used toefficiently and hemostatically ablate prostate tissue. These lasers arepumped by CW krypton arc lamps and produce a constant train ofQ-switched pulses at 25 kHz. The high Q-Switch frequency makes thetissue effects indistinguishable from CW lasers of the same averagepower. The 532 nm light from these lasers is selectively absorbed byblood leading to good hemostasis. When ablative power densities areused, a superficial layer of denatured prostate tissue less than 1 mm isleft behind. This thin layer of denatured tissue is thin enough that theimmediate post surgical swelling associated with other treatmentmodalities is greatly reduced. This reduced swelling leads to shortcatheterization times and less dysuria. At high powers, 532 nm lasersinduce a superficial char layer (an absorptive, denatured layer) thatstrongly absorbs the laser light and greatly improves the ablationefficiency. The problem with the existing 532 nm lasers used to date isthat they are large, expensive, inefficient, and have a highlymulti-mode output beam that makes them inefficient for ablating prostatetissue. Furthermore, residual coagulation of tissue due to the procedureremains significant using the techniques known in the prior art, asdiscussed below.

High power densities are required for rapid and efficient vaporizationof prostate tissue. The difficulty of achieving higher average outputpower densities is that when high input powers are supplied to the laserelement from an excitation source such as an arclamp a large amount ofheat is generated in the lasing element. This heat induces variousdeleterious effects in the lasing element. In particular the temperaturedifference between the coolant and the hot lasing element generates athermally induced graded index lens that decreases the beam quality ofthe laser and causes the laser to operate with more transverse opticalmodes than it would otherwise.

The M² parameter is a well established convention for defining the beamquality of a laser and is discussed in pages 480-482 of Orazio Sveltoand David C. Hanna, Principles of Lasers, Plenum Press, New York, 1998,which is incorporated herein by reference. The beam quality measures thedegree to which the intensity distribution is Gaussian. The quantity M²is sometimes called inverse beam quality rather than beam quality but inthis application it will be referred to as beam quality. M² is definedas${{M_{x}^{2} \equiv \frac{\left( {\sigma_{x}\sigma_{f}} \right)_{NG}}{\left( {\sigma_{x}\sigma_{f}} \right)_{G}}} = {4{\pi\left( {\sigma_{x}\sigma_{f}} \right)}_{NG}}},$

where π refers to the number 3.14 . . . , σ is used to represent thespot size, the subscripts x and f represent the spatial and frequencydomains along the x-axis, respectively, and the subscripts G and NGsignify Gaussian and non-Gaussian, respectively. The x-axis istransverse to the direction of propagation of the beam. The beam qualityin any direction transverse to the beam may be essentially the same.Therefore the subscript x is dropped from the M² elsewhere in thespecification. The beam widths or σs are determined based on thestandard deviation of the position, where the squared deviation of eachposition is weighted by the intensity at that point. The beam width inthe frequency domain σ_(f) is the beam width of the beam after beingFourier transformed.

The formula usually used for calculating the angular divergence, θ, of abeam of light of wavelength λ is strictly valid only for a beam having aGaussian intensity distribution. The concept of beam quality facilitatesthe derivation of the angular divergence, θ, for the beam with anon-Gaussian intensity distribution, according to$\theta = {{M^{2}\left( \frac{2\quad\lambda}{\pi\quad\sigma_{x}} \right)}.}$

For example, a TEM00 laser beam has a high beam quality with an M² of 1,whereas by comparison, high power surgical lasers operate with M² valuesgreater than 100.

The Applicants have recognized that high power lasers typically have anM²>144. The larger number of modes makes M² larger and makes itdifficult to focus the light into small, low numerical aperture fibersand reduces the ability to project high power density light onto tissue.As a result, the vaporization efficiency of CW arclamp pumped 532 nmlasers on prostate tissue is significantly reduced.

Other aspects and advantages of the present invention can be seen onreview of the drawings, the detailed description and the claims, whichfollow.

SUMMARY OF THE INVENTION

Photoselective vaporization of tissue, such as the prostate fortreatment of BPH, is based upon applying a high intensity radiation toprostate tissue using a radiation that is highly absorptive in thetissue, while being absorbed only to a negligible degree by water orother irrigant during the operation, at power densities such that themajority of the energy is converted to vaporization of the tissuewithout significant residual coagulation of adjacent tissue. Unlikeprior art techniques for treatment of BPH, the procedure may beconducted under local anesthesia, and patients are usually able to gohome a couple of hours after the procedure. The procedure results infewer side effects than prior art techniques, including lower incidenceof dysuria and hemouria. Patients may be treated without requiringpost-operative catherization of the urethra.

According to one embodiment of the invention, a method for treating BPHcomprises the steps of providing a solid-state laser having a laserelement positioned to receive pump radiation from an excitation source;in some cases modulating the source to cause the laser to emit pulsedlaser light; and delivering the laser light to targeted tissue. Varioussolid-state lasers may be used for this purpose, including (withoutlimitation), a Q-switched laser using a frequency doubling crystal suchas potassium-titanyl-phosphate (KTP), pumped using a diode array, an arclamp or a flash lamp. While Q-switching induces short, “micro-pulses,” a“macro-pulse” duration of the laser light is preferably in the range of0.1 to 500 milliseconds, induced by for example modulating the pumpenergy with the desired macro-pulse length. The wavelength of the laserlight is preferably between 200 and 1000 nm. The laser light ispreferably delivered to the targeted prostate tissue through an opticalfiber terminating at or near a distal end in a side-firing probe.However the side-firing probe is not essential.

Operation of the solid-state laser in a “macro-pulsed” mode is moreefficient in inducing rapid tissue ablation than a CW laser of the sameaverage power. This is in part because the macro-pulsing is moreefficient in inducing “char” formation, a mild carbonization in whichthe tissue typically darkens slightly but does not necessarily turncompletely black. Although char formation is not essential to efficientrapid ablation it is helpful because the darkened tissue is better atabsorbing light. The macro-pulsed laser is also more efficient and hashigher beam quality, with M² values typically less than 144, than acontinuous wave laser with same average output power.

According to a second embodiment of the invention, a method for treatingsoft tissue comprises the steps of providing a solid-state laser havinga laser element positioned to receive pump radiation from a pumpradiation source; modulating the pump radiation source to cause thelaser element to emit laser light having a pulse duration of between 0.1milliseconds and 500 milliseconds and an output power exceeding 20watts; and delivering the laser light to targeted tissue.

According to a third embodiment of the invention, a method for treatingBPH comprises the steps of providing a solid-state laser having a laserelement positioned to receive pump radiation from a pump radiationsource; Q-switching the laser to generate a quasi-continuous wave (CW)beam having an output power exceeding 60 watts; and, delivering the beamto targeted prostate tissue.

According to a fourth embodiment of the invention, a method for treatingBPH comprises the steps of providing a solid-state laser having a laserelement positioned to receive pump radiation from a pump radiationsource such as a laser diode; Q-switching the laser to generate aquasi-continuous wave (CW) beam having an output power exceeding 20watts with an M² less than 144; and delivering the beam to prostatetissue.

It has been recognized that as more and more laser energy is consumed byvaporization of the tissue, the amount of laser energy leading toresidual tissue coagulation gets smaller, i.e. the amount of residualcoagulation drops, and the side effects attendant to the residual injurycaused by the surgery drop dramatically. Thus, the extent of the zone ofthermal damage characterized by tissue coagulation left after theprocedure gets smaller with increasing volumetric power density, whilethe rate of vaporization increases. Substantial and surprisingimprovement in results is achieved. It has been recognized thatincreasing the volumetric power density absorbed in the tissue to beablated, has the result of decreasing the extent of residual injury ofthe surrounding tissue. This recognition leads to the use of higherpower laser systems, with greater levels of irradiance at the treatmentarea on the tissue, while achieving the lower levels of adverse sideeffects and a quicker operation times.

Although the invention can be generalized other types of tissue, oneembodiment of the invention provides a method for photoselectivevaporization of prostate tissue. According to this embodiment, themethod includes delivering laser radiation to the treatment area on thetissue, via an optical fiber for example, wherein the laser radiationhas a wavelength and irradiance in the treatment area on the surface ofthe tissue sufficient because vaporization of a substantially greatervolume of tissue than a volume of residual coagulated tissue caused bythe laser radiation. In one embodiment, the laser radiation is generatedusing a neodymium doped solid-state laser, including optics producing asecond or higher harmonic output with greater than 60 watts averageoutput power, and for example 80 watts average output power, or more.The laser radiation is coupled into an optical fiber adapted to directlaser radiation from the fiber to the treatment area on the surface ofthe tissue. For the treatment of prostate, the fiber optic is insertedvia transurethral cystoscope, including lumens for delivering irrigantsto the treatment area, and for direct visualization during thetreatment.

In other embodiments, the delivered laser radiation has a wavelength ina range of about 200 nm to about 650 nm, and has an average irradiancein the treatment area greater than about 10 kilowatts/cm², in a spotsize of at least 0.05 mm². More preferably, the irradiance is greaterthan about 20 kilowatts/cm², and even more preferably greater than about30 kilowatts/cm². The spot size in preferred systems is for example lessthan about 0.8 mm².

Accordingly, in one embodiment, the second harmonic output of theneodymium dope solid-state laser is applied using a side firing tip onthe optical fiber. The side firing tip, which causes a diverging beam tobe directed out of the optical fiber, is placed close to the tissue,within about 1 mm from the side of the side firing tip to contacting theside of the tip. Close placement increases the irradiance delivered tothe treatment area so that higher irradiance is available withsolid-state lasers generating a 60 to 80 watts average output power.

According to the present invention, the efficiency of the vaporizationand the reduction of injury to residual tissue are sufficient that theprocedure may be carried out while applying only local anesthetic duringthe delivery of laser energy, and throughout the procedure. For example,a procedure according to the present invention includes applyingintraurethral topical anesthesia such as lidocaine, either aperiprostatic block or a perirectal block, oral and/or intravenous drugssuch as Fentanyl or Demerol, chilled irrigant, and irrigant containinganesthesia.

Furthermore, embodiments of the invention include the delivery of thelaser energy using a Q-switched, solid-state laser which producesmicro-pulses in combination with applying pump power to the laser mediumin a sequence a pulses so that output radiation is produced inmacro-pulses having a peak power of greater than 200 watts, and morepreferably about 240 watts or greater. The peak irradiance in thetreatment area during the pulses is thereby substantially increased, andpreferably greater than 50 kilowatts/cm², and as much as 90kilowatts/cm² in some embodiments of the invention.

Other aspects and advantages of the present invention can be seen onreview the figures, the detailed description, and the claims whichfollow.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts a laser system for implementing the tissue ablationmethods of the invention;

FIG. 2 depicts a side-firing probe for use with the system of FIG. 1;

FIG. 3 depicts an exemplary output waveform of the FIG. 1 laser when thelaser is operated in a macro-pulsed mode; and

FIG. 4 depicts an exemplary output waveform of the FIG. 1 laser when thelaser is operated in a quasi-CW mode.

FIG. 5 is a block diagram of a laser system adaptable for use accordingto the present invention.

FIG. 6 is a block diagram of an alternative laser system adaptable foruse according to the present invention.

FIG. 7 is a diagram of a transurethral cystoscope, adaptable for useaccording to the present invention.

FIG. 8 illustrates absorption depth in prostate tissue for 532 nm light.

FIG. 9 illustrates absorption depth in prostate tissue for 1064 nmlight.

FIG. 10 is a diagram of a beam path from an end view of a side firingtip, according to one embodiment of the present invention.

FIG. 11 is a diagram of a beam path from a side view of the side firingtip of FIG. 10, according to one embodiment of the present invention.

FIG. 12 is a heuristic diagram illustrating operation of the presentinvention.

DETAILED DESCRIPTION

FIG. 1 is a block diagram depicting an exemplary laser system 100 whichmay be employed for implementing the present invention. Laser system 100includes a solid-state laser 102, which is used to generate laser lightfor delivery through optical fiber 106 to target tissue 104. As will bediscussed in further detail herein below, laser 102 is capable of beingoperated in a “macro-pulsed” mode, wherein the laser light is emitted asmacro-pulses having relatively long pulse durations.

Laser 102 more specifically comprises a laser element assembly 110, pumpsource 112, and frequency doubling crystal 122. In the preferredembodiment, laser element 110 outputs 1064 nm light which is focusedinto frequency doubling crystal 122 to create 532 nm light. According toone implementation, laser element assembly 110 may be neodymium dopedYAG (Nd:YAG)crystal, which emits light having a wavelength of 1064 nm(infrared light) when excited by pump source 112. Laser element 110 mayalternatively be fabricated from any suitable material whereintransition and lanthanide metal ions are disposed within a crystallinehost (such as YAG, Lithium Yttrium Fluoride, Sapphire, Alexandrite,Spinel, Yttrium Orthoaluminate, Potassium Gadolinium Tungstate, YttriumOrthovandate, or Lanthanum Scandium Borate). Laser element 110 ispositioned proximal to pump source 112 and may be arranged in parallelrelation therewith, although other geometries and configurations may beemployed.

Pump source 112 may be any device or apparatus operable to excite laserelement assembly 110. Non-limiting examples of devices which may be usedas pump source 112, include: arc lamps, flashlamps, and laser diodes.

A Q-switch 114 disposed within laser 102 may be operated in a repetitivemode to cause a train of micro-pulses to be generated by laser 102.Typically the micro-pulses are less than 1 microsecond in durationseparated by about 40 microseconds, creating a quasi-continuous wavetrain. Q-switch 114 is preferably of the acousto-optic type, but mayalternatively comprise a mechanical device such as a rotating prism oraperture, an electro-optical device, or a saturable absorber.

Laser 102 is provided with a control system 116 for controlling andoperating laser 102. Control system 116 will typically include a controlprocessor which receives input from user controls (including but notlimited to a beam on/off control, a beam power control, and a pulseduration control) and processes the input to accordingly generate outputsignals for adjusting characteristics of the output beam to match theuser inputted values or conditions. With respect to pulse durationadjustment, control system 116 applies an output signal to a powersupply (not shown) driving pump source 112 which modulates the energysupplied thereto, in turn controlling the pulse duration of the outputbeam.

Although FIG. 1 shows an internal frequency doubled laser, it is only byway of example. The infrared light can be internally or externallyfrequency doubled using non-linear crystals such as KTP, LithiumTriborate (LBO), or Beta Barium Borate (BBO) to produce second harmonic532 nm green light, and higher harmonics. The frequency doubled, 532 nmwavelength and the shorter wavelength higher harmonic beams are betterabsorbed by the tissue, and promote more efficient tissue ablation.

In one preferred embodiment the resonant cavity control system is thatdescribed in U.S. Pat. No. 5,151,909, which is incorporated by referenceas if fully set forth herein.

Laser 102 further includes an output port couplable to optical fiber106. Output port 118 directs the light generated by laser 102 intooptical fiber 106 for delivery to tissue 104. Mirrors 124, 126, 128, and130 direct light from the lasing element 110 to the frequency doublingcrystal 122, in addition to forming the resonant cavity of the laser.Mirrors 124, 126, 128, and 130 are configured for focusing the light toform an image just in front of the frequency doubling crystal 122 on theside closer to mirror 130, and to compensate for thermal lensing in thelasing element. Although mirrors 124, 126, 128, and 130 are illustratedas flat and parallel to the walls of the laser, typically the focusingis achieved by curving and/or angling the mirrors. Alternativelytransmissive optical elements could be used to focus the light andcompensate for the thermal imaging. Mirrors 124, 128 and 130 reflectboth the wavelength of light produced by the lasing element (e.g. 1064nm) and the wavelength of the frequency doubled light (e.g. 532 nm).Mirror 126 only reflects the light originating from the lasing element110 (e.g. 1064 nm) but is transparent to the frequency doubled light(e.g. 532 nm), forming an output window. Higher harmonic outputs mayalso be generated from the 1064 nm line, or other line amplified in thelaser, including third and fourth harmonics, for shorter wavelengths.Other laser systems may be used, including but not limited to Sapphirelasers, diode lasers, and dye lasers, which are adapted to provide theoutput power and wavelengths described herein, including wavelengths inthe ranges from 200 nm to 1000 nm and from 1100 nm to 1800 nm, forexample.

While a bare fiber may be utilized for certain procedures, optical fiber106 preferably terminates in a tip 140 having optical elements forshaping and/or orienting the beam emitted by optical fiber 106 so as tooptimize the tissue ablation process.

FIG. 2 depicts a side-firing probe tip 200, which may be used as tip 140(FIG. 1). The tip 140 is treated to deflect light sideways. Someexamples of methods for deflecting the light sideways are to include alight scattering material in the tip 140 and/or to place a reflectiveelement in the tip 140. The reflective element could be angled at 45°,for example; to deflect the light at 90° with respect to the axis of thefiber 106. Side-firing probe tip 200 includes an optically transparentsleeve 202 having a transparent window 204 (which may be constructed asa cutout in the wall of sleeve 202 through which the beam is emitted ina direction transverse to the optical axis of fiber 106.) An acceptablerange of angles in which to deflect the light beam is between about 40to 120 degrees with respect to the axis of the fiber. The preferredembodiments use an angle of either 70 or 100. The angle of 80° ispreferred from the standpoint of the ease in manufacturing the tip 200and the angle of 90° is preferred from the standpoint of the ease inaiming the side firing light.

In a typical mode of operation, optical fiber 106 is held within anendoscope such as a cystoscope or similar instrument that allows theclinician to precisely position the distal end of the optical fiberadjacent to the targeted tissue. The endoscope also has channels forsupplying and removing an irrigant solution to and from the tissue. Inaddition, light and image guides are also included for illuminating andimaging the tissue so that the clinician may direct the laser light andassess the progress and efficacy of the ablation procedure.

FIG. 3 illustrates an exemplary output waveform applied to tissue 104when laser 102 is operated in the macro-pulsed mode. Each macro-pulse302 is defined by a train of Q-switched micro-pulses 304. While arelatively small number of micro-pulses 302 are depicted for purposes ofclarity, an actual macro-pulse may comprise hundreds or thousands ofcomponent micro-pulses 304. In the preferred embodiment there arebetween 2 and 12,200 micro-pulses per macro-pulse.

An arc lamp, for example, when used as the pump source 112, is kept at alow power level between pulses that are preferably just enough tomaintain the arc. These low pump powers are below the lasing thresholdof the laser; as a consequence, there is no laser output betweenmacro-pulses.

As mentioned above, the pulse duration or width D (FIG. 3) of the outputbeam is governed by the modulation of pump source 112, and morespecifically by the period during which the pump source 112 ismaintained in an “on” or high-power condition. In other words, thelonger the pump source 112 is maintained in an on condition, the longerthe pulse width. Typically, laser 102 will be capable of deliveringpulses 302 having pulse durations D in the range of 1 to 20 milliseconds(2 to 490 micro-pulses) or 1 to 50 milliseconds (2 to 1,220micro-pulses) and average output powers preferably exceeding 60 wattsand preferably up to 100 or 200 watts. The ratio of D to the period ofthe macro-pulses defines the duty cycle, which is typically between 10and 50%.

In accordance with one embodiment of the invention, a laser system 100of the foregoing description is employed to treat BPH by ablatingtargeted prostate tissue 104. The clinician may utilize an endoscope orsimilar instrument to guide the distal end and tip 140 of optical fiber106 into alignment with the targeted prostate tissue 104.

Laser system 100 is then operated in the macro-pulsed mode so that laser102 generates laser light having the pulsed waveform depicted in FIG. 3and delivers it through optical fiber 106 to tissue 104.

It is known that irradiation of prostate tissue 104 may initially causetissue heating resulting in the formation of a char layer. This charlayer is highly optically absorptive in the wavelengths emitted by laser102, which thereby facilitates efficient absorption of the laser lightand resultant ablation of tissue 104. However, the formation of the charlayer is not essential for efficient ablation. Prior art techniques fortreatment of BPH by laser ablation (such as the technique described byKuntzman et al. in “High-Power (60-Watt) Potassium-Titanyl-PhosphateLaser Vaporization Prostatectomy in Living Canines and in Human andCanine Cadavers,” Urology, Vol. 49, No. 5 (1997)) utilized a quasi-CWlaser to irradiate the prostate. Although such lasers do producemoderately high average powers, they have a large number of transversemodes and as such, produce highly divergent light when focused intosmall fiberoptics. This leads to less than optimal power densities whenthe laser light is directed at tissue. As a consequence, these lasersare not particularly efficient at inducing formation of a char layer,and ablation rates are relatively slow, significantly lengtheningprocedure times. Further, since formation of the char layer takes placeat relatively low rates, undesirable thermal damage to deeper tissuelayers may occur. In contrast, it has been found that a macro-pulsedbeam, such as that generated by laser 102, promotes rapid formation of achar layer even at moderate output energy levels, thereby helping toaccelerate ablation rates and reducing procedure time.

The macropulsing can also increase efficiency because the thresholdvoltage required for lasing while macropulsing (the operating threshold)is lower than the initial threshold voltage for lasing (cold threshold).

Macropulsing is also more efficient for producing green light becausethe conversion of infrared light to frequency doubled light increases asthe square of the infrared light intensity. The higher peak powers ofthe macro-pulsed infrared light leads to higher second harmonicconversion efficiency. For example, at any given time, the input powerand output power of a frequency-doubled laser using KTP are relatedaccording toPo=A(Pi)²,

Where A is an experimentally determined positive constant. This equationrelates the peak input power to the peak output power. However, theaverage input power and output power for a duty cycle of k percent aregiven by<Pi>=k(Pi) and<Po>=k(Po)=kA(Pi)² =A(<Pi>)² /k,where the brackets “< >” indicate an average value of the enclosedquantity. Thus, decreasing the duty cycle from 100% to 50% (i.e.reducing k from 1 to 0.5) while simultaneously doubling the peak inputpower Pi results in no change to the average input power <Pi> and adoubling of the average output power <Po>. Pulse modulating ormacropulsing using Q-switching, for example, enables reaching higheraverage output powers with less thermal lensing due to the lower inputpower.

Additionally, it is possible that the frequency doubling crystal hasnonlinearly increasing output power as a function of the input power. Inother words the second derivative of the output power with respect tothe input power may be positive, in which case the rate of increase ofthe output power increases with increasing input power. Specifically, insuch a case the functional dependence of the instantaneous or peakoutput power, Po, on the instantaneous or peak input power, Pi, is suchthatd ²(Po)/d(Pi)²>0.When this is true, and Po is an increasing function of Pi, the higherpeak input power results in a more efficient laser because ratio of theoutput to input power increases.

Pump source modulation of the laser can produce high peak powermacro-pulses and affect the efficiency of the average power output.Macro-pulse in excess of a steady state power can substantially improvethe initiation of the vaporization of prostate tissue. The higher peakpower of the macro-pulse rapidly initiates charring which in turn servesas an additional chromophore for the incident energy and enhances thevaporization rate. A 30% macro-pulse duty cycle is sufficient toincrease the average power output of an arc lamp pumped laser to greaterthan 80 watts. Additionally the pump modulation generates macro-pulsewith pulse powers greater than 240 watts.

By way of a non-limiting example, prostate tissue 104 may be efficientlyand rapidly ablated when laser 102 is operated at an output power of 80to 100 watts, a pulse duration of 1-50 milliseconds, and a wavelength of532 nm.

In accordance with a second method embodiment of the invention, lasersystem 100 may be utilized to ablate other types of tissue 104.Treatment of tissue 104 is performed in a manner substantially identicalto the technique for treating BPH disclosed above. The clinician mayutilize an endoscope or similar instrument to guide the distal end andtip 140 of optical fiber 106 into alignment with the prostate tissue104. Laser system 100 is then operated in the macro-pulsed mode so thatlaser light having the pulsed waveform depicted in FIG. 3 is generatedby laser 102 and delivered through optical fiber 106 to tissue 104. Toachieve adequate results, laser system 100 is adjusted to emit a beamhaving a pulse duration between 0.1 and 500 milliseconds, and an outputpower of at least 20 watts. Upon vaporization of the required volume oftissue 104, (which may be assessed via an imaging channel contained inthe endoscope), the output beam of laser 102 is turned off.

In a third method embodiment of the invention, treatment of BPH iseffected by operating laser 102 in a quasi-CW mode at an output powergreater than 60 watts. The increased denaturization of the tissue isdramatic with increases in power, suggesting a threshold effect. Asdepicted in FIG. 4, laser 102 generates a continuous train of Q-switchedmicro-pulses 400 when operated in quasi-CW mode. The laser light is thendelivered via optical fiber 106 to targeted tissue 104. Operation in aquasi-CW mode at powers above 60 watts facilitates formation of char andconsequent rapid ablation rates, whereas operation in a quasi-CW mode atpowers below 60 watts forms char more slowly and causes more thermaldamage to underling tissue.

A fourth embodiment of this invention is to produce a high power, highbeam quality laser that can project high power density laser light ontotissue. To do this the number of transverse optical modes supported bythe resonator needs to be kept as low as possible.

Small M² and high average powers can be achieved by reducing the degreeof thermal lensing in the laser element. Using laser diodes as theexcitation source is one effective way of greatly reducing both the sizeof the lasing element and the thermal gradient responsible for creatingthe thermal lens. The reason for this is that while 2-10% of the lightproduced from a flashlamp or arc lamp is converted into useful laserlight 30-60% of the light emitted from laser diodes can be converted tolaser light. Since the energy that is not converted to laser light isconverted into heat, laser diodes deposit significantly less heat in thelasing element and as a consequence create a less powerful thermal lens.In this manner laser diodes can be used to pump crystalline laserelements or fiber lasers to produce high beam quality lasers. Slab andwaveguide lasers that can be pumped by laser diodes, arc lamps, orflashlamps are another method of creating low M² lasers. This is becausethe thermal gradient produced in slab lasers is linear across the thindimension of the slab and not radially dependent in contrast to atypical cylindrical lasing element. The linear thermal gradient does notproduce a thermal lens resulting in low M² values.

For example, as a result of the low M² some embodiments of thisinvention are capable producing laser light that upon exiting a flat endof a fiber having a diameter of 600 μm has a divergence of 15.3° orlower; 15° or lower; 10° or lower; or 5° or lower, and the power densitycan be 13,400 watts per cm², or greater.

FIG. 5 shows a block diagram of a preferred laser system according tothe present invention. In FIG. 5, a laser resonator is defined by endmirror 10, turning mirrors 12 and 14, and end mirror 16. All of thesemirrors are high reflecting (greater than 99.8%) at the 1064 nm line. Anoptical path 24 is defined by these mirrors. A gain medium 18 comprisinga Nd:YAG rod is mounted along the optical path within a lamp housing 29.An arc lamp 28 is also mounted within the housing and supplies pumppower to the gain medium in response to current generated in powersupply 30. Also in the optical path 24 is a Q-switch 20 between the lamphousing 29 and the turning mirror 12. A non-linear crystal 22 is mountedbetween the turning mirror 14 and the back mirror 16. This non-linearcrystal is preferably a KTP crystal aligned for frequency doubling togenerate a 532 nm beam. Mirrors 16 and 14 are highly reflective at 532nm, while mirror 12 is transmissive and operates as an output couplerfor the 532 nm beam.

Thus, the laser resonator is designed for resonating at a firstfrequency, i.e., 1064 nm along the Z-shaped optical path 24. A secondfrequency derived from the 1064 nm beam is generated in the KTP crystal22. This beam travels along the path 26 a and is extracted from theresonator to supply an output beam along path 26 b.

The output beam along path 26 b passes through a controllable attenuator36, a beam splitter 38, which supplies a portion of the output beam to asurgical detector 40, and a component group 42 as described in moredetail below. The attenuator, detector, and component group are allcoupled to a data processing system 34, across lines 34 j, 34 k, and 34p.

The Q-switch 20 is controlled by Q-switch driver 21, which is, in turn,coupled to data processor 34 across line 34 i. In the preferred system,the Q-switch is an acoustic-optic Q-switch.

Similarly, the power supply 30 generates an electrical power signal forcontrolling the arc lamp 28. This power signal is controlled by the dataprocessor 34 across line 34 h and by drive circuitry 32 across line 32a. Drive circuitry 32 a is controlled by the data processor across lines34 a through 34 g. A sensor 57 is coupled with the data processor tosense an environmental condition, such as temperature or humidity, thataffects operation of the laser system. A modem 56 is connected to thedata processor 34, providing an interface for remote access to memory inthe data processor. Finally, a control panel 35, by which a user cansupply input signals and parameters, is provided. This control panel 35is connected to the data processor 34 across line 34 n.

In alternative systems, the non-linear crystal may be mounted outsidethe resonant cavity of the resonator. Also, it may be used forextracting outputs other than the second harmonic, such assum-of-frequency derivation or the like.

The wavelength used according to the present invention for BPH treatmentshould be strongly absorbed in the prostate tissue to help initiate andmaintain tissue vaporization without creating deep tissue heating. Thewavelength also must be minimally absorbed by the irrigant it usedduring the procedure, typically water. The 532 nm light produced by thesystem of FIG. 5, is both strongly absorbed in oxyhemoglobin and weaklyabsorbed in water. Oxyhemoglobin is readily present in prostate tissueand serves as an efficient chromophore for 532 nm light. Thedifferential in absorption coefficients between oxyhemoglobin and waterat 532 nm is approximately 5 orders of magnitude (10⁵). In otherembodiments, wavelengths in the range from 200 nm-650 nm are used, whichhave strong oxyhemoglobin absorption and relatively weak waterabsorption (>10²X). In yet other embodiments, wavelengths in the rangefrom 200 nm to 650 nm range are used, which have strong oxyhemoglobinabsorption and relatively weak water absorption (>10X).

Of course, as shown in FIG. 6, in which like components have the samereference numerals as in FIG. 5, alternative pump power sources, such aslaser diode arrays, other lasers for longitudinal pumping, and others,can be used as suits the needs of a particular gain medium andapplication of the laser system. Representative laser diodes includelaser diodes providing output in the range of 805 to 820 nm inwavelength with an input power to the array of pumping diodes in therange of 300 to 500 Watts. The laser diodes used for pump energy areoperated in a modulated macro-pulse mode, or in a continuous mode, assuits a particular implementation.

The laser systems shown in FIGS. 5 and 6 can be modified by removingboth the Q-switch and the external surgical attenuator. The Q-switch andsurgical attenuator are removed because the modulated pump powerprovides a great deal of flexibility in controlling the output power ofthe laser not attainable using a Q-switch. The data processing systemcan be programmed to maintain a constant thermal load in the lasersystem while varying the peak pump power widely. Thus, the peak currentand duty cycle of the pump power source can be adjusted in such a way tokeep the average pump power constant, but the second harmonic powerduring the ready and work modes adjusted by selecting the peak currentand duty cycle. Although it may be necessary to use attenuators in thebeamline during the ready mode in order to extract an aim beam, suchattenuators may well be eliminated for the work mode. The average powerdoes not have to be constant, rather it can be maintained at levelswhich keep thermal focusing of the gain medium within the range ofstability of the resonator.

A representative laser system adapted for delivery of energy asdescribed above, comprises an 80 watt average power, 532 nm outputwavelength, solid state, intra-cavity frequency doubled Nd:YAG laser. Toobtain optimal efficiency, an arc lamp pump source is modulated at aperiod of 4.5 ms with a 16 ms duty cycle, generating 285 watts peakmacro-pulse power. An intra-cavity acousto-optic AO Q-switch is used tofurther modulate the energy at a period of 40 kHz with a 450 usmicro-pulse. The laser energy is coupled to a side firing fiber opticdelivery device for delivery to prostate tissue.

The laser system uses a combination touch screen and control knob userinterface to assist the operator in setting up the surgical parameters,including power levels and pulse sequence specifications. The averagepower setting is prominently displayed on the screen. Parameteradjustments are made by first activating (touching) the desiredparameter box on the screen and then turning the knob. The laser systemuses a secure card key interface to enable the laser. The system istransportable. The system offers convenient storage and a fiber deliverydevice pole.

An example of an endoscope, in particular a transurethral cystoscope,for use with the present invention is shown in FIG. 7. The cystoscopehas a distal end 200 and a proximal end 201. The distal end 200 includesa tongue member 202 for pushing tissue away from a treatment area in theregion 204. Laser radiation is directed sideways from opening 206 in tothe region 204, by a side firing fiber optic component. Water or salinesolution is delivered and removed from the treatment area via lumens inthe probe. A viewing optic is also placed in the opening 206, by whichthe surgeon is able to view the treatment area during the procedure. Onthe proximal end 201 of the cystoscope, an input port 203 and an outputport 205 for flow of the irrigant is provided. Also, a fiber port 207 isused for insertion and removal of the fiber optic delivering laserradiation to the treatment area. A light source connector 209 is usedfor supplying light to the treatment area for visualization. A telescope211, which can be coupled to a video camera, or looked into directly, isalso included on the representative cystoscope.

The vaporization of prostate tissue using oxyhemoglobin as the primarychromophore is related to the incident power density, or irradiance,which can be expressed in Watts/cm². The overall rate of prostate tissuevaporization is a function of the spot size, absorption depth, and thepower density. A large spot with high power density, and rapidabsorption is ideal to rapidly vaporize tissue. A high power lightsource is required to achieve a large spot, high power density treatmentbeam. Peak laser power, average laser power, beam quality, deliverydevice design and delivery device placement all affect the efficiency ofvaporization. A treatment beam 28.5 Kw/cm² average irradiance with a85.5 Kw/cm² peak irradiance macro-pulse , with a spot size between about0.2 and 0.5 mm², rapidly vaporizes tissue.

The BPH treatment procedure can be outlined as follows for oneembodiment of the invention, using a laser system as described abovewith reference to FIG. 5.

-   -   Equipment/Set-up        -   21-24 french continuous flow cystoscope        -   Laser        -   Side firing probe        -   Filter (eye safety filter for the monocular or the video            camera)        -   Sterile water        -   Cystoscope eye piece or video system for direct            visualization    -   Anesthesia        -   Any of the following or combinations of the following:            General, spinal nerve block, topical, periprostatic block,            perirectal block, pudental block & intervenous drugs        -   This procedure does not require general or spinal anesthesia    -   Technique/Process        -   Prep patient following standard protocol for cystoscopic            procedures        -   Administer anesthesia        -   Dilate the urethra        -   Insert cystoscope        -   Insert side firing delivery device        -   Begin flow of sterile water        -   Cystoscopy to assess gland        -   Position fiber near tissue to be removed and active the            laser        -   Use a sweeping motion to vaporize desired tissue            -   Continue vaporization until the capsule is reached            -   Monitor vaporization efficiency, remove and clean fiber                as required        -   Debulk desired lobes, median & lateral lobes        -   Fill bladder with water, remove cystoscope, observe            discharge        -   If necessary insert foley catheter

The rapid vaporization with thin coagulation zone contribute to thehemostasis during the procedure. Because of minimal thermal damage toexisting tissue, there is a low incidence of side effects, making suchsymptoms as Dysuria, Incontinence and Impotence which often occur inprior art techniques, very unlikely.

Further, the procedure causes minimal bleeding. Great outcomes areachieved for patients suffering BPH, including improved urine flow rate,improved post-void residual, and improved symptom scores on BPH tests.The procedure often achieves immediate obstruction relief, and postoperative catheterization is not always required.

A typical photoselective vaporization of the prostate PVP procedure willuse the following steps:

A. At the investigator discretion, van buren sound the urethra in astandard fashion prior to insertion of the continuous flow cystoscope.

B. Subjects will be administered general, spinal or local (prostaticblock & oral and topical anesthesia) anesthesia at the discretion of theSurgeon. In some embodiments, the procedure is performed without the useof general anesthesia or spinal nerve blocks, using only localanesthesia such as any combination of intraurethral topical anesthesiasuch as lidocaine, either a periprostatic block or a perirectal block,oral and/or intravenous drugs such as Fentanyl or Demerol, chilledirrigant, and irrigant containing anesthesia.

C. Vaporization will be performed with the Laserscope ADD (AngledDelivery Device) fiber, which is a 600 um bare fiber with a quartzcapsule over the 70 degree lateral deflecting quartz element and a spotdiameter of 1.2 mm at 2 mm.

D. The laser fiber will be introduced through the lumen of a standard 22Fr continuous flow laser cystoscope, and sterile water will be used asthe irrigant.

E. KTP laser energy will be generated by a high power 532 nm lasercapable of delivering 80 W of KTP laser power to tissue.

F. Lasing will be performed under direct visualization using a free beamtechnique, holding the fiber 1-2 mm away from the tissue and vaporizingthe lateral lobes beginning at the bladder neck.

G. The visible laser beam will be slowly moved along the length andbreadth of the lateral lobe as the tissue is vaporized. The laser willbe carefully directed toward the apical tissue making sure to protectthe external sphincter.

H. Both lateral lobes will be vaporized evenly to the level of thecapsular fibers.

I. The median lobe will be vaporized evenly to the level of thetransverse fibers of the vesicle neck. If the median lobe is too large,then it should be partially vaporized before ablation of the laterallobes to facilitate the movement of the scope and irrigation, and thenthe remainder will be flattened out later during the procedure.

J. The procedure should preserve the distal crista urethralis and theverumontanum.

K. The end point of the procedure should be judged by the size andappearance of the large transurethral resection-like cavity and by thediminished efficacy of the vaporization effect at the prostatic capsule.The median and lateral hypertrophic tissue must be vaporized to thelevel of the transverse fibers and any lingering loose fibers should beremoved prior to completion of the treatment.

L. Rarely will arterial bleeders be encountered, however, if an arterialbleeder is encountered, then coagulate the vessel at a distance ofapproximately 3 to 4 mm.

M. The cystoscope is removed and, if necessary, a foley catheter isinserted at the physician's discretion.

N. This is an outpatient procedure and subjects will be released fromthe hospital as outpatients per the discretion of the Surgeon.

FIGS. 8 and 9 illustrate the different optical penetration depths of the532 nm wavelength and 1064 nm wavelength used in prior art procedures.See, S. L. Jacques. Laser-tissue interaction. Photochemical,photothermal, and photomechanical. Surg. Clin N. Am. 1992;72(3):531-558.The optical penetration depth of the 1064 wavelength beam from Nd:YAGlaser beam is about 10 mm, which is 13 times larger than the penetrationdepth of the second harmonic 532 wavelength laser beam, which is about0.8 mm. As a result, the 1064 laser power is spread out over a muchlarger tissue volume than the power of the KTP laser. In case of the1064 laser as shown in FIG. 9, the temperature at the tissue surfacebarely reaches 100° C. Therefore, only a small portion of tissue getsvaporized. But a huge volume of tissue gets coagulated (see spacebetween 100° C. and 60° C. isotherm). The consequences are the formationof edema in a huge volume of coagulated tissue, swelling of theprostate, and the patient going into retention with catheterizationtimes of several weeks.

The 532 laser beam, in contrast, is substantially completely absorbedwithin less than about 1 mm of the surface of prostatic tissue. Thelaser power is confined to a very small tissue volume. The highvolumetric power density results in a fast heating of the tissue andefficient tissue vaporization. Volumetric power density delivered totissue is a function of the absorption depth, irradiance in Watts/cm²and spot size on the surface of the tissue. The coagulation zone is verythin because of the small optical penetration depth of the 532wavelength, and because substantially all of the radiation is convertedto vaporization rather than residual heat.

Other wavelengths which are substantially completely absorbed withinless than about 1 mm of the surface of the prostatic tissue includewavelengths less than about 650 nm, for example between about 200 nm and650 nm.

FIGS. 10 and 11 illustrate a profile of a beam delivered to tissue usingone representative side firing optical fiber, to show spot size as afunction of distance from the side of the optical fiber. FIG. 10 is anend view, showing a fiber 600, cladding 601 on the fiber, an air space602, and a tip 603 through which the beam is directed by a reflectingface on the fiber. The cross-section of the beam is represented by thecrossing lines 604 and 605. As shown, the beam has a width in thisdimension of about 0.35 mm at 1 mm from the side of the tip 603. Atabout 2 mm from the side of the tip 603, the width is about 1 mm. Atabout 3 mm distance from the side of the tip 603, the beam width isabout 2.2 mm.

FIG. 11 is a side view, with like components given the same referencenumbers. The beam width in this dimension is represented by lines 606and 607. As shown, the beam has a width in this dimension of about 0.7mm at 1 mm from the side of the tip 603. At about 2 mm from the side ofthe tip 603, the width is about 1 mm. At about 3 mm distance from theside of the tip 603, the beam width is about 1.5 mm.

Thus, the spot size at 1 mm from the side of the tip is definedbasically by an elipse having a major axis of 0.7 mm, and a minor axisof 0.35 mm. The area of the spot at 1 mm is around 0.2 mm². At 2 mm fromthe side, the area of the spot is about 0.8 mm².

For rapid procedures, according to the present invention, the spot sizeshould be large enough that the operator can remove tissue at areasonable rate, and see the results of a single pass of the spot over aregion of tissue. If the spot size is too small, the rate of theoperation is too slow. Also, if the spot size is too big, then theprocedure is difficult to control precisely. A preferred spot size isless than about 1 mm, and more particularly between about 0.8 mm² andabout 0.05 mm². Other apparatus may be used for delivery of the beamwith the desired spot size, including embodiments without divergingbeams, and embodiments with converging beams.

The irradiance of the beam at 1 mm from the side of the tip for an 80 Waverage power laser as described above is about 30 kiloWatts/cm².According to the present invention, it is desirable to provide awavelength between about 650 and 200 nm, with a spot size on the surfaceof the tissue less than about 0.8 mm², and preferably greater than about0.05 mm², with an irradiance greater than about 10 kiloWatts/cm², andmore preferably greater than 20 kiloWatts/cm², and even more preferably30 kiloWatts/cm² or higher.

FIG. 12 shows, heuristically, how vaporization rate and coagulation ratedepend on the volumetric power density. The vaporization rate (in mm/s)is defined as tissue depth that is vaporized per time interval. Thecoagulation rate (in mm/s) is defined as the depth of residualcoagulated tissue that remains after a certain time of vaporization.

Below a certain volumetric power density, referred to as a “vaporizationthreshold” in FIG. 12, no tissue gets vaporized. All laser energy staysinside the tissue. Tissue coagulation occurs where the tissuetemperature rises above approximately 60° C. As the volumetric powerdensity is increased a bigger and bigger tissue volume gets coagulated.

At the vaporization threshold, vaporization starts. Above thevaporization threshold the vaporization rate can be considered toincrease linearly with the volumetric power density for the purpose ofunderstanding the present invention, and as described by a steady statemodel for continuous wave laser tissue ablation, known by those familiarwith the art of laser-tissue interaction.

As more and more laser energy is consumed by vaporization of the tissue,the amount of laser energy leading to residual tissue coagulation getssmaller, i.e. the amount of residual coagulation drops. Thus, extent ofthe zone of thermal damage characterized by tissue coagulation leftafter the procedure gets smaller with increasing volumetric powerdensity, while the rate of vaporization increases. Substantial andsurprising improvement in results is achieved.

Publications about visual laser ablation of the prostate (VLAP) that isperformed with an Nd:YAG laser at 1064 nm have shown that this type oflaser is not able to vaporize a significant amount of tissue. Histologystudies have shown that the 1064 nm laser induces deep coagulation inthe tissue that results in edema and delayed tissue sloughing. Thiseffect was described by Kuntzman, et al., High-power potassium titanylphosphate laser vaporization prostatectomy. Mayo Clin Proc1998:73:798-801. Thus, in the heuristic diagram of FIG. 12, the VLAPprocedure is believed to lie around point 650, barely above thevaporization threshold. Also, prior art technologies using 532 nm withspot sizes on the order of 1 mm² with average output power of 60 Watts,are believed to lie, heuristically, around point 651 in the FIG. 12.Kuntzman et al present results for the coagulation depth of a 60 Wcontinuous wave 532 nm laser, with suggested operation at a distance of2 mm from the side of the tip, yielding less than 5 kiloWatts/cm²irradiance.

As the laser power is further increased to 80 W, and the side firingprobe is placed less than 1 mm from the tissue for a small spot size,the ablation rate further increases and the coagulation rate furtherdrops, so that the procedure lies heuristically at point 652 in FIG. 12.

An 80 Watt KTP laser can used to easily reach irradiance levels thatvaporize substantially more tissue than is left as residual coagulationafter the procedure. More precisely, the vaporization rate issubstantially higher than the coagulation rate as given by thedefinition above, using high irradiance levels that are easily achievedwith higher power lasers.

While the invention has been particularly shown and described withreference to preferred embodiments thereof, it will be understood bythose skilled in the art, that various changes in form and details maybe made therein without departing from the spirit and scope of theinvention, as defined by the appended claims.

1. A method for photoselective vaporization of tissue, comprising:delivering laser radiation to a treatment area on the tissue, the laserradiation having a wavelength and having irradiance in the treatmentarea sufficient to cause vaporization of a substantially greater volumeof tissue than a volume of residual coagulated tissue caused by thelaser radiation.
 2. A method for treating benign prostatic hypertrophy(BPH), comprising the steps of: providing a solid-state laser emittinglight with a wavelength of 200 to 1000 nm having a laser elementpositioned to receive pump radiation from a pump radiation source;modulating the pump radiation source to cause the laser element to emitlaser light having pulse duration between 0.1 and 500 milliseconds andpulse frequencies between 1 and 500 Hz; and delivering the laser lightto targeted tissue.
 3. A method for treating benign prostatichypertrophy (BPH), comprising the steps of: providing a solid-statelaser having laser element positioned to receive pump radiation from apump radiation source; Q-switching the laser to generate aquasi-continuous beam having an output power exceeding 60 watts, anddelivering the beam to targeted prostate tissue.
 4. The method of claim3 wherein the output power is 80 watts or greater.
 5. A methodcomprising: hemostatically irradiating soft tissue with coherentradiation from a laser having a pump radiation source that does notproduce enough heat in a lasing element to create a significant amountof thermal lensing in the lasing element; and a power and wavelengthsuch that the radiation is preferentially absorbed by the tissue ratherthan a fluid.
 6. An apparatus for photoselective vaporization of tissue,comprising: a laser producing laser radiation; an endoscope, includingan optical fiber coupled to the laser, adapted to direct laser radiationfrom the fiber, and a flow of irrigant to a treatment area on a surfaceof the tissue; laser and optical fiber being adapted to deliver thelaser radiation at a wavelength and irradiance in the treatment areasufficient to cause vaporization of a substantially greater volume oftissue than a volume of residual coagulated tissue caused by the laserradiation.